1. Field of the Invention
The present invention relates generally to imaging devices for nuclear medicine, and more specifically relates to gamma-ray or scintillation cameras and methods of obtaining images from radiation data acquired by such cameras by measuring the position of a radiation event in not only the x-y plane but also in the z plane.
2. Background and Prior Art
A gamma-ray or scintillation camera as utilized in nuclear medicine is a well known device. The original scintillation camera or "Anger camera" (named after the inventor) is described in U.S. Pat. No. 3,011,057. The Anger camera uses a scintillation crystal, such as a NaI(Tl) crystal, which absorbs incident gamma rays from the object under study and interacts with the gamma ray to produce light events. An array of photomultiplier tubes is placed adjacent to the crystal in order to detect and amplify these light events so as calculate the spatial location and energy level of the incident gamma ray to produce a two dimensional image of the object which then may be displayed on a CRT or printed as a hard copy.
When a nuclear medical image is being acquired, a radioisotope has been introduced into the body as a radiopharmaceutical having an affinity for certain parts or organs of the body, and the diagnostician is interested in the distribution of that radiopharmaceutical within the body or organ under evaluation. It is therefore desirable that the image accurately represent the spatial distribution of the radiation emitted from the body. When radioactive nuclei decay, a gamma-ray or high energy X-ray is emitted from the location of the decay. The gamma-rays (or X-rays) travel in a straight line until they are either scattered or absorbed. If a gamma-ray is absorbed in the scintillation crystal of the camera and detected as a light event without having undergone an intervening scattering event, then the location at which the gamma-ray was detected represents the actual location of the decay, and hence part of the distribution of the radioisotope. Such an event is considered a "good" detected event and is used to form an accurate picture of the radioisotope distribution within the body. However, if the gamma-ray scatters within the body at some point between its emission from the location of decay and its detection in the scintillation crystal, then the location at which the scattered gamma-ray is detected does not represent the location from which the gamma-ray was emitted, and thus the inclusion of such an event in the image will falsely indicate the presence of radioisotope where, in actuality, there may not have been any radioisotope. Such an event is known as a "bad" event.
The phenomenon by which a gamma-ray collides with an electron (of an atom of the body, for example), loses some of its energy and changes its direction of travel is known as Compton scattering. Because the scattered gamma-ray energy is lower than the energy of the unscattered gamma-ray, it is the energy of a gamma-ray event that is used to discriminate among detected gamma-ray events so as to include only unscattered gamma-ray events in the image being acquired. When a single gamma-ray is absorbed in the scintillation crystal, a fraction of the deposited energy is emitted as scintillation photons which have wavelengths within the visible spectrum. Because the scintillation photons are emitted isotropically from the point of absorption, only a small amount of the emitted photons reach the photomultiplier tubes (PMTs). The fraction of the total amount of photons reaching the photomultiplier tubes that produces an electrical signal in any one photomultiplier tube is dependent on the position of that photomultiplier tube relative to the location of the light event, local variations in physical properties of the crystal, reflective surfaces, other transparent media such as lightpipes, and the interfaces between all of these materials and the boundaries of the detector. Additionally, the probability that a scintillation photon entering a photomultiplier tube will be converted into an electrical pulse is dependent on local variations in the photocathode of the photomultiplier tube. This probability is known as the quantum efficiency of the photocathode. The quantum efficiency is highly dependent on the thickness and composition of the photocathode, and is thus variable from PMT to PMT as well as locally within a PMT.
An alternative to the sodium iodide crystal-PMT type gamma ray camera is disclosed in U.S. Pat. No. 5,171,998, issued Dec. 15, 1992 to two of the present inventors. The disclosure of the '998 patent is incorporated by reference herein in its entirety. The '998 patent describes a gamma ray imaging detector having a scintillating crystal made of CsI(Tl) (thallium-doped cesium iodide) instead of NaI(Tl), and also including an array of photodiodes arranged on one side of the scintillating crystal opposite the side which receives incident gamma rays, for receiving scintillation photons produced by the interaction of the incident gamma ray with the crystal. A light reflective surface overlies the surface of the crystal opposite the photodiode array.
As described in the '998 patent, the CsI(Tl) crystal exhibits a higher scintillation photon yield than NaI(Tl), albeit in the 580 nanometer wavelength range, as opposed to the 415 nanometer range for NaI (in which PMTs are more sensitive to incident photons). However, when combined with the use of low noise, low capacitance photodiode detectors, which have a higher quantum efficiency than conventional photomultiplier tubes in the 580 nanometer range, the alternative gamma ray imaging detector disclosed in the '998 patent achieves an energy resolution significantly improved over the conventional PMT-type detector.
Conventionally, the measurement of the spatial location of a detected gamma ray absorption event in the scintillating crystal has been limited to a two-dimensional point in the X,Y plane of the crystal. However, because the number of scintillation photons reaching each detector element (either a PMT or a photodiode) is dependent on the solid angle subtended by the area of that detector element to the point of the gamma ray absorption within the crystal, the amount of scintillation photons received by each detector is also a function of the depth of interaction (DOI) of the incident gamma ray within the crystal, i.e., along the Z axis of the crystal. The DOI is an important parameter when applied to imaging detector geometries in which the directions from which incident gamma rays impinge upon the crystal are not all substantially normal to the crystal surface. If incident gamma rays intersect the crystal from directions not normal to the crystal, the unknown depth of interaction of those gamma rays within the crystal will result in an additional uncertainty in the measured position of the interaction because of the parallax effect, if only a two dimensional (i.e., X,Y) spatial location is calculated for such an absorption event. A detailed explanation of the importance of and the problems associated with the DOI is provided in "Maximum Likelihood Positioning in the Scintillation Camera Using Depth of Interaction," D. Gagnon et al., IEEE Transactions on Medical Imaging, Vol. 12, No. 1, March 1993, pp. 101-107.
Consequently, there exists a need in the art for a gamma ray imaging detector which is capable of measuring not only the spatial location of gamma ray absorption events in a scintillating crystal in terms of X,Y coordinates, but which is also capable of measuring the depth of interaction of the event within the crystal in the Z axis direction, so as to improve the accuracy and performance of the imaging function.